Closed microfluidic network for strain sensing embedded in a contact lens to monitor intraocular pressure

ABSTRACT

A microfluidic strain sensing device for monitoring intraocular pressure. The device has a contact lens and a closed microfluidic network embedded with the contact lens. The network has a volume that is sensitive to an applied strain. The network distinguishes: (i) a gas reservoir containing a gas, (ii) a liquid reservoir containing a liquid that changes volume when the strain is applied, and (iii) a sensing channel able to hold the liquid within the sensing channel. The sensing channel connects the gas reservoir on one end and connects the liquid reservoir on another end. The sensing channel establishes a liquid-gas equilibrium pressure to interface and equilibrium within the sensing channel, which would fluidically change as a response to radius of curvature variations on a cornea, or as a response to mechanical stretching and release of the cornea. The liquid-gas equilibrium pressure interface and equilibrium are used for measuring the intraocular pressure.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims priority from U.S. Provisional PatentApplication 62/556,366 filed Sep. 9, 2017, which is incorporated hereinby reference.

FIELD OF THE INVENTION

This invention relates to devices, systems and methods to monitorintraocular pressure. In particular, the invention relates tomicrofluidic network design for strain sensors, which work based onmechanical amplification of the volume of the microfluidic channels, tomonitor intraocular pressure.

BACKGROUND OF THE INVENTION

Glaucoma is a neurodegenerative disease that causes irreversible damageto eye's optic nerve, and hence, loss of vision. Continuous andlong-term monitoring of intraocular pressure (IOP) is critical formanagement of glaucoma.

IOP reduction is the only known way of slowing and/or stopping theprogression of glaucoma. It is estimated that for every 1 mmHg of IOPreduction, the risk of nerve damage is reduced by 11%. Drug therapy iscommonly used to reduce IOP, but there are important challenges thatneed to be addressed to improve effectiveness of glaucoma treatment.Most importantly, nearly 50% of patients stop using medication after sixmonths for various reasons. Continuous, long-term IOP monitoring thathas the capability to measure drug efficacy could help patients staycompliant and help physicians with the management of glaucoma. Moreover,in recent years, diurnal variations in IOP is established as anotherrisk factor for glaucoma, which increased the importance of continuousmeasurements even further.

Current technologies available for IOP measurements are either notcontinuous (Goldmann Applanation Tonometry), or continuous but temporary(Sensimed Triggerfish) or continuous but invasive (implantable sensors).The self-tonometry devices (e.g. Icare) can provide long-term data andit is noninvasive but still uncomfortable for the patient to a levelthat it may require topical anesthetics. Moreover, the results obtainedby self-tonometry are found to be user dependent.

Approaches for telemetric continuous IOP measurements have beendeveloped and tested in animal models. Amongst these approaches, contactlens-based monitoring techniques are attractive because they arenon-invasive. One contact lens system (Sensimed AG's Triggerfish)measures minute changes in corneal curvature by a contact lens equippedwith an electrical strain sensor, an antenna and a microchip that areused for processing and transmitting signal wirelessly. The technologyrequires the patient to wear a receiver on the waist for datatransmission and power transfer. Because of the thick silicone contactlens (central thickness 580 μm), it is not as comfortable as daily usedcontact lenses; a mild to moderate adverse reactions are reported in upto 80% of patients. The requirement of a trained personnel, anddiscomfort and high cost associated with this contact lens platformprecludes its usage in long term monitoring applications but only allowstests for a single 24-hour period. For this reason, Triggerfish found tobe more suitable for determining changes in IOP in a daily scale.However, IOP change as a response to a drug is in the time scale ofweeks. Similarly, IOP change in response to certain lifestylemodifications will also be in longer than 24-hour time scale. Therefore,there is a need for a continuous wear contact lens sensor that canmonitor IOP variations in long-term for determining the drug efficacyand for decreasing the number of doctor visits that a patient needs tomake for routine IOP measurements.

Other examples of contact lens sensors are based on measurement ofelectrical resistance, inductance and capacitance changes in response topressure induced strain. In these examples, sensor response is typicallydetected remotely by measurement of the resonant frequency changes usingan external reader coil or by Bluetooth connectivity. The electricalmeasurements require conductive components inside the lenses, which aretypically not transparent and not air permeable.

Recently, Kim et al. used graphene-Ag-nanowire to address the electrodetransparency issue (J. Kim et al., “Wearable smart sensor systemsintegrated on soft contact lenses for wireless ocular diagnostics,”Nature Communications, vol. 8, April 2017, Art. no. 14997). Firstcondition of a contact lens with long-term usage capability is high airpermeability to prevent hypoxia. Unfavorably, the conductive componentsneeded by electrical sensors are impermeable to gases. Metals have 8-10orders of magnitude lower gas permeability compared to soft materialsand this cause mild adverse reactions in human trials when electricalsensing-based contact lenses are used even for a single 24 hours. Theother condition for long term usage is comfort, which is achieved bymaking high water content and thin (<200 micrometer) contact lenses. Theelectrical sensing methods are sensitive to the hydration level of thecontact lens. Therefore, the contact lens electrical sensors are made ofsilicone, which have very low water content, instead of the standardsilicone/hydrogel materials. This reduces the comfort of the contactlens. There are three main reasons for sensitivity to hydration level.First, swelling of the hydrogel due to hydration induces a strain, andtherefore, it is a source of error in measurement. Second, the frictionbetween contact lens and cornea can be sensitive to hydration level,hence influences the sensitivity. Finally, the electrical components areaffected from the humidity, and therefore, should be isolated by usingsealant materials such as parylene-c.

The present invention advances the art and provides technology tomeasure IOP eliminating at least some of the current problems orconcerns.

SUMMARY OF THE INVENTION

The invention pertains to a strain sensor using microfluidic principlesintegrated with a contact lens for IOP measurements. The materials usedin the invention are low-cost, transparent, air permeable, and flexible.A method is provided to embed the microfluidic strain sensor in asilicone contact lens. The microfluidic contact lens sensor (miLenS)allows patients to measure their own IOP to better manage the glaucoma.

The microfluidic contact lens sensor is capable of measuring the IOPfluctuations due to internal (i.e. metabolism, blinking and saccadic eyemovements) as well as external factors (i.e. drugs, diet, lifestyleetc.) during the lifetime of the patient. The measurements will be doneat the discretion of the patient (or automatically) where readout willbe realized by a smartphone camera (or by a wearable camera forautomated measurements). This allows for at-home monitoring andcontinuous data recording. The data then will be sent directly to amedical provider's database, which allows patients and physicians tomonitor IOP variations. Aspects of our technology are listed as follows:

-   1) The miLenS will be built using a hybrid material system where the    narrow microfluidic sensing region (as low as 0.1 mm wide ring at    the periphery of miLenS) is embedded in a silicone or    silicone/hydrogel contact lens material. The microfluidic sensing    channels will be made out of transparent, soft, and oleophobic    materials. The sensing material will be 6-10 orders of magnitude    more air permeable compared to electronic components.-   2) The microfluidic sensing technique has no actively controlled    components and only works based on the principles of fluid physics.    The miLenS is free of all electrical components (powerless). It is a    low-cost device. Additionally, this provides easier usability by    eliminating the cumbersome peripheral components (e.g. antenna,    microchip etc.) for data transmission, reception and recording,    which are needed in wearable electronic sensors.-   3) The sensor will be sensitive to strain and responds to corneal    radius of curvature changes but has low sensitivity to forces    applied directly by the eyelid or due to hydration of the contact    lens materials. The sensor we designed has low stiffness in lateral    direction (i.e. the microfluidic device is thin and has low elastic    modulus) and high stiffness in radial direction (i.e. the    microfluidic network channels have small width), which will make it    insensitive to external forces (e.g. blinking, rubbing of the eye).-   4) The miLenS enables readout with a smartphone camera and an    optical adaptor. This will provide measurements at discrete time    points. In one variation, a wearable camera that can track the    sensor response can also be utilized for continuous and automated    measurements.-   5) Continuous data recorded with existing technologies show that IOP    fluctuates around 5-15 mmHg day-to-day and hour-to-hour, and 15-40    mmHg second-to-second. The microfluidic network circuitry we have    designed has the capability to filter out large fluctuations that    occurs in short time scales due to blood pressure or muscle    contractions. In this case, the sensor actually acts as a fluidic    low-pass filter, which only responds to changes that occur in    minutes or slower. In a similar manner, the fluidic components can    be designed to register only the rapid IOP changes. A sensor that    can measure events happening in different time scales can make    better estimation of the true IOP based on the corneal radius of    curvature measurements.

The microfluidic strain sensor embedded contact lens is convenient touse and has continuous measurement capability. It requires minimaltraining to take measurements therefore, will be used as a device forhome-medicine. These will enable clinical studies where recording oflong-term IOP data on large patient populations is needed. Continuousrecording of IOP and its analysis will improve our understanding ofneurodegenerative diseases and their relation to pressure. Additionally,it will be useful for improving the efficiency and efficacy of drugsthat are used for glaucoma treatment. Therefore, miLenS technologyoffers a promising healthcare technology for better personalized care ofglaucoma patients. These advantages listed above will potentially enablethe patient to use the sensor permanently and without the trainedpersonnel assistance.

In one embodiment, the present invention provides a microfluidic strainsensing device for monitoring intraocular pressure changes. The closedmicrofluidic network is transparent and/or oleophobic. The microfluidicstrain sensing device has a contact lens and a dosed microfluidicnetwork embedded with the contact lens. The contact lens is a siliconecontact lens, a hydrogel contact lens or a combination thereof. Thecontact lens has no actively controlled components or electricalcomponents.

The closed microfluidic network has a volume that is sensitive to anaxial strain. The closed microfluidic network distinguishes: (i) a gasreservoir containing a gas, (ii) a liquid reservoir containing a liquidthat changes volume when the strain is induced, and (iii) a sensingchannel able to hold the liquid within the sensing channel. The sensingchannel connects the gas reservoir on one end and connects the liquidreservoir on another end. The sensing channel establishes a liquid-gasequilibrium pressure interface and equilibrium within the sensingchannel, which would fluidically change as a response to radius ofcurvature variations on a cornea, or as a response to mechanicalstretching and release of the cornea. The liquid-gas equilibriumpressure interface and equilibrium are used for measuring theintraocular pressure.

The liquid reservoir forms at least one ring and wherein the airreservoir is positioned inside or outside the at least one ring. In eachcase, the liquid reservoir volume is highly sensitive to tangentialforces on the eye relative to radial forces on an eye wearing thecontact lens. The liquid reservoir has a high stiffness in radialdirection and/or smaller channel width relative to the stiffness intangential direction and/or a microfluidic channel wall thicknessresulting in the liquid reservoir becoming insensitive to externalforces.

In one example, the liquid reservoir has one or more chambers. Thesechambers could have concentric rings. These chambers could also haveconcentric rings that are connected to each other at one or morelocations. These chambers could also have concentric rings where thesensitivity increases as the number of concentric rings increases.

In one example, the surface of the liquid reservoir could be patterned.The surface of the liquid reservoir ceiling could have a convex shapeand the convex shape could be curved towards the reservoir channelfloor.

The sensing channel has a strain sensitivity of about 4.5 mm interfacemovement per about 1% strain applied to the liquid reservoir. In oneexample, the sensing channel has an inner diameter of about 1-10 mm. Inanother example, the sensing channel has an inner diameter 5-12 mm witha cross sectional area of 10⁻¹¹-10⁻⁸ m².

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 shows according to an exemplary embodiment of the invention aworkflow of the miLens device based on pressure monitoring.

FIG. 2A shows according to an exemplary embodiment of the invention animage of a sensor which is only 100 micrometers thick. The small dropson each side are Norland Optical Adhesive (NOA) used to seal the sensorand can be made less than 20 micrometers thick.

FIG. 2B shows according to an exemplary embodiment of the invention animage of a sensor after embedding the sensor into the contact lens (300micrometers final thickness).

FIG. 3 shows according to an exemplary embodiment of the invention a topview of a closed system sensor with multiple ring liquid reservoirembedded in a contact lens.

FIG. 4 shows according to an exemplary embodiment of the invention aside view of a multiple chamber liquid reservoir sensor A) versus asingle chamber liquid reservoir sensor B) and their respective behaviorwhen the sensors are stretched under tangential forces as shown in A*)and B*). 410-A and 410-B show possible stretch points under stretchingof the sensor. The sensor has to be made from a soft material decreasingthe stiffness in both directions. The sensor has to be thin. FIG. 4illustrates this: basically, the microfluidic channel ceiling thicknesst1 and floor thickness t2 has to be small (<20 μm). This also reducesthe stiffness in both directions. The reservoir ring width w has to besmall (<100 μm). This does not affect the tangential stiffness butincreases the radial stiffness of the microfluidic channel and is key inincreasing the sensor performance.

FIG. 5 shows according to an exemplary embodiment of the invention topviews of a single ring liquid reservoir versus a three rings reservoir.The circled region for the three rings shows the rings zoomed in.

FIG. 6 shows according to an exemplary embodiment of the inventionpressure response of three different sensor types; 1, 2, and 5 reservoirrings. Ring height, width and separation is 100 micrometers. The slopevalues are the sensitivity and shown under corresponding curve in mm permmHg unit. For each curve, the average and standard deviation of atleast 3 measurements are used.

FIG. 7 shows according to an exemplary embodiment of the invention asensitivity dependence on the number of reservoir rings for threedifferent ring widths. The multiple data points for some of the ringnumbers are obtained using sensors fabricated at different times withthe same parameters; fluctuations in sensitivity values are in result offabrication variances. The sensitivity depends linearly on the number ofrings with 50 and 100 micrometer widths, but no significantly affectedfrom it for 200 micrometer width.

FIG. 8 shows according to an exemplary embodiment of the invention aside view of the placement of the miLenS on a cornea and position of theliquid reservoirs. The insets show the close-up look of the liquidreservoirs and the forces acting on them; inset a) shows a single wideliquid reservoir compressed under radial force, and inset b) shows aseries of concentric circles as liquid reservoir not compressed underthe same force. As this figure is shown in grey scale, for colors toreflect the effect of the forces, please refer to U.S. ProvisionalPatent Application 62/556,366 filed Sep. 9, 2017, which is incorporatedherein by reference.

FIG. 9 shows according to an exemplary embodiment of the inventionsensitivity dependence on the height for three different ring widths.The multiple data points for some of the heights are obtained usingsensors fabricated at different times with same parameters; fluctuationsin sensitivity values are the result of fabrication variances. Thesensitivity depends linearly on the reservoir height. The red datapoints 910 indicate thicker chips (300 micrometers) and they show 50%reduced sensitivity compared to thinner (150 micrometers) counterparts920. As this figure is shown in grey scale, for colors, please refer toU.S. Provisional Patent Application 62/556,366 filed Sep. 9, 2017, whichis incorporated herein by reference.

FIG. 10 shows according to an exemplary embodiment of the invention anauxetic contact lens sensor and close-up view of the liquid reservoircross section. Unlike a sensor with a rectangular channel as shown inFIG. 8, this channel has a curved top layer. This top layer getsflattened when tangential force is applied as shown according to ourdata and Comsol simulations. As this figure is shown in grey scale, forcolors, please refer to U.S. Provisional Patent Application 62/556,366filed Sep. 9, 2017, which is incorporated herein by reference.

FIG. 11 shows according to an exemplary embodiment of the invention asensor with a reservoir ceiling patterned with circular and linearconvex shapes.

FIG. 12 shows according to an exemplary embodiment of the invention amicroscope image of the sensor to the left with a linearly patternedliquid reservoir ceiling. To the right is shown a comparison of measuredsensitivity for a flat ceiling sensor verses a curved ceiling (auxetic)device 29 micrometer/mmHg and 77 micrometer/mmHg, respectively.

FIG. 13 shows according to an exemplary embodiment of the invention amethod of fabricating the sensor. A refers to UV treatment. B refers toplasma treatment (PDMS). C refers to treatment APTES. 1 refers to glassslide, 2 refers to NOA65 (uncured), 3 refers to PDMS, 4 refers to NOA65(cured). Step 1 is NOA65 sandwiched in between two PDMS coated glassslides and UV cured to create 20 micrometer films. This is repeatedtwice. Step 2 is NOA65 dropped on the mold and 20 micrometer films fromstep 2 is plasma treated. Step 3 is two layers from step 2 aresandwiched together and UV cured. Step 4 is a 70 micrometers layer fromstep 3 is plasma treated. The 20 micrometers layer from step 1 is plasmatreated and APTES treated. Step 5 is two layers from step 4 aresandwiched together. As this figure is shown in grey scale, for colors,please refer to U.S. Provisional Patent Application 62/556,366 filedSep. 9, 2017, which is incorporated herein by reference.

FIG. 14 shows according to an exemplary embodiment of the invention amethod of embedding the sensor into a contact lens. B refers to plasmatreatment (PDMS). C refers to treatment APTES. D refers to cure (heat)treatment. 5 refers to the hemispherical mold for contact lensfabrication, 6 refers to the sensor, 7 refers to the top layer of thecontact lens. Step 6 is PDMS poured on a contact lens mold. Then curedat 80 degrees Celsius, plasma and APTES treated. Sensor bottom surfaceis plasma treated. Step 7 is sensor bottom surface placed on the PDMScoated contact lens mold. The sensor reservoirs are filled with workingliquid and sealed. Step 8 is more PDMS poured on the sensor and cured atroom temperature. Step 9 is contact lens peeled off the surface of themold. As this figure is shown in grey scale, for colors, please refer toU.S. Provisional Patent Application 62/556,366 filed Sep. 9, 2017, whichis incorporated herein by reference.

FIG. 15 shows according to an exemplary embodiment of the inventionfabrication steps of a ceiling layer of the auxetic microfluidic sensor.

FIG. 16 shows according to an exemplary embodiment of the invention astrain sensor for biomechanics of cancer cells. In the bottom channel, astrain sensor is placed meanwhile cells are seeded in the top channel.

FIG. 17 shows according to an exemplary embodiment of the invention atop view and a side view of a contact lens and the location of theshapes. Besides star shapes in the top view and side view other exampleshapes can also be provided. The combinations of these shapes can alsobe used.

FIG. 18 shows according to an exemplary embodiment of the inventionComsol results where 50 micrometers high and 50 micrometers widechannels provide close to the optimal sensitivity while maintaining thindevices. Star shape in FIG. 18 shows the optimal geometrical parametersfor maximum volume change (i.e., sensitivity) while maintaining thedevice thin.

DETAILED DESCRIPTION

IOP measurement devices reported so far do not consider thedirectionality of the forces acting on the sensors. For example,capacitance measurement-based sensor by Chen et al. (G.-Z. Chen, I.-S.Chan, L. K. K. Leung, and D. C. C. Lam, “Soft wearable contact lenssensor for continuous intraocular pressure monitoring,” MedicalEngineering & Physics, vol. 36, no. 9, pp. 1134-1139, September 2014),responds to the radial forces applied on the lens such as due toblinking. An ideal contact lens sensor should only be sensitive to thestrain applied in result of the radius change of the cornea, but shouldnot be affected by the forces applied perpendicularly on the lens (i.e.radial forces). In consideration of this, we have used COMSOLsimulations and experimental measurements to develop a strain sensor,which is more sensitive to tangential forces than the radial forces onthe eye. Embodiments of the invention are based on microfluidic sensingfor IOP measurements and such desired strain sensor force response.

FIG. 1 shows an example of a workflow of the IOP self-measurementtechnique. The miLenS is distinct from other sensors because patientswill be able to place and remove it by themselves similar to a regularcontact lens. As IOP fluctuates, radius of corneal curvature changes(each 1 mmHg change in IOP causes 4 μm change in radius of curvature).In this technology, the fluidic level in the microfluidic sensingchannel of the sensor will change as a response to radius of curvaturevariations on the cornea. The sensor response will be detected with asmartphone camera equipped with an optical adaptor and then converted topressure value by a smartphone app. It will eliminate the security andhealth concerns related to radio frequency or Bluetooth data transfermethods. We have demonstrated an IOP detection limit of 1 mmHg onenucleated porcine eyes, which is sufficient for IOP monitoringapplications.

Microfluidic circuits, analogous to electronic circuits, can function aslow or high pass filters (electrical resistance and capacitance arereplaced by fluidic resistance (R) and the compliance (C) ofcompressible materials, respectively). The RC value will determine thetime constant of the sensor response. Sensors with large RC values willnot respond to fast changes but will be sensitive to slowly varyingdiurnal variations. Sensors with small RC values will have thecapability to detect the effects of blinking and ocular pulsation.

In one exemplary embodiment, the microfluidic strain sensor (FIG. 2A) isintegrated into a PDMS contact lens (FIG. 2B) for wearable sensingapplications. Referring to FIGS. 3-4, sensor 300 with sensor material302, the sensor 300 is embedded in contact lens 310 distinguishes aliquid reservoir 320 (amplifies the displaced liquid volume and show inthis example as liquid reservoir rings), a gas reservoir 330 and asensing channel 340 connected to the liquid reservoir 320 on one end andto a gas reservoir 330 on the other. First, liquid reservoir 320 isfilled with a working liquid such as oil using capillary action and thensealed. This creates a stable gas/liquid interface 350 in the sensingchannel 340 and forms a closed microfluidic network. The IOPfluctuations change the corneal radius of curvature; for every 1 mmHgincrease in IOP, corneal radius of curvature increases 4 μm. Thisincreases the liquid reservoir volume due to the strain applied on theliquid reservoir elastic walls. The increased reservoir volume creates avacuum and shifts the gas/liquid position 350 in the sensing channel 340towards the liquid reservoir 320. As the sensing channel cross sectionarea is reduced, the linear liquid displacement required to accommodatethe reservoir volume change increases, hence the sensitivity improves.

FIG. 5 shows the top view of two example designs—single ring 510 versusthree rings 520 for the liquid reservoir—of the microfluidic strainsensor. Increasing the vertical wall surface area of the liquidreservoir increase the sensitivity of the sensor to changes in IOP. Thiswas tested in two ways; i) increasing the number of walls ii) increasingthe height of the channel walls. First, we designed and fabricatedsensors with multiple liquid reservoir rings as shown by for example520, thus increasing the total wall surface area. The sensitivityresults for different number of rings are presented in FIGS. 6-7. Wefound that increasing the number of walls by adding more rings,increased the sensitivity of the device in a linear manner. On thecontrary, the width of the reservoir did not have a significant effecton the sensitivity. This phenomenon is a direct result of the interplaybetween tangential strain and radial force induced collapses as shown inFIG. 8. To test the effect of the reservoir wall height we built threetypes of sensors (50, 100 and 330 μm height) and compared theirsensitivity. As shown in FIG. 9, as the reservoir height doubled, thesensitivity is also doubled. When we increased the reservoir height to330 μm the sensitivity also increased by a factor of three (shown onlyfor 200 μm width), proving the effect of vertical wall height. FIG. 9further shows the effect of sensor stiffness. When a 150 μm thick sensoris compared to a 300 μm thick one (shown by 100T and 330T), the thickersensors have ˜50% lower sensitivity.

In summary, we experimentally scanned a large parameter range tounderstand and optimize the sensor performance. We have fabricatedsensors with varying number of reservoir rings (1-5), ring widths(w=50-500 μm), reservoir heights (50, 100, 330 μm) and chip thicknesses(130 μm, 300 μm) as well as different Young's moduli ˜1 MPa (PDMS) vs˜10 MPa (NOA 65) and ˜100 MPa (NOA 61). The results of these sensitivitytests have demonstrated that; i) The increased liquid reservoir heightincreases the sensitivity. ii) We are able to improve the sensitivity byadding more reservoir rings to the design as needed (e.g. depending onthe required continuous wear contact lens properties). iii) Thestiffness (Young's modulus (E)×chip thickness (t)/width (w)) does notalter the sensitivity significantly; however, it needs to be optimizedin view of other factors such as comfort and lens/cornea mechanicinteractions.

Auxetic Metamaterials for Microfluidic Strain Sensing

In another version of the sensor, the microfluidic channel networkheight increases in response to the applied tangential strain 1010. Thevolume increase is achieved by Poisson ratio modification throughlithographical patterning of elastomeric sensor. FIG. 10 shows, via across-section of the contact lens sensor, the working principle of theauxetic metamaterials for strain sensing. The ceiling of themicrofluidic channel has a convex shape, i.e. curved towards the channelinterior, as shown. This is achieved by patterning the ceiling film witheither circular or linear patterns as shown in FIG. 11. Although theseare the only patterns we tested other patterns can be used to get thesame effect. When the tangential force is applied (i.e. due to IOPchanges), as shown in FIG. 10, the ceiling gets deformed outward becauseof the convex ceiling, as opposed to the collapses observed when flatceiling is used. This deformation towards the front face of the sensorcauses a channel height increase, hence amplification in liquidreservoir volume expansion, according to our COMSOL simulations as shownin U.S. Provisional Patent Application 62/556,366 filed Sep. 9, 2017(FIG. 14 therein), which is incorporated herein by reference. Thisamplification increases the sensitivity of the sensor.

FIG. 12 on the left shows the image of the liquid reservoir on anauxetic sensor with a linear pattern of convex structures on theceiling. FIG. 12 on the right shows the experimental sensitivitycomparison between flat and curved (auxetic) devices. The sensitivityincrease is 2.5-fold.

Microfluidic mechanical metamaterials that are biocompatible andelectronics-free enabled fabrication of highly sensitive and reliablestrain sensors. The tangential strain-sensing method we developed isspecific to IOP as demonstrated by our experiments. We have used thisapproach to monitor IOP in porcine eyes and demonstrated 1-mmHgdetection limit (corresponds to 0.05% strain) and reliability forseveral weeks. The microfluidic strain sensor can measure the strain ofthe eye due to the shape changes in response to IOP in the clinicallyrelevant range.

Manufacturing

We built the sensors using photolithography and soft lithographytechniques. First, polydimethylsiloxane (PDMS) soft molds werefabricated. As a sensor material, polyurethane based Norland OpticalAdhesive 65 (NOA65) was chosen due to its transparency, flexibility,oleophobicity and biocompatibility. Then, thin NOA65 films with therequired features were made and bonded together to make sensors as shownin FIG. 13. For the purposes of this invention, we developed specificfabrication methods to build extremely thin (˜100 μm) microfluidicdevices. The gas permeability of polyurethane used in our devices is 6-8orders of magnitude lower than metals used in wearable electronics.

We have first cut the strain sensor into desired shape and embedded aflat 100 μm strain sensor (FIG. 2A) into a PDMS contact lens. Althoughwe have built our sensors flat they can also be built curved if curvedmolds were used. We have developed a fabrication protocol where we canbuild contact lenses with 8-15 mm radius of curvature and 10-14 mmradius as shown in FIG. 2B. We have used dome shaped plastic molds wherewe poured PDMS on them to obtain the 10-100 μm silicone film at thedesired radius of curvature, we bonded our sensor on the silicone filmby (3-Aminopropyl) triethoxysilane (APTES) chemistry. Then, poured moresilicone to fully embed the sensor in silicone. The details are shown inFIG. 14. Finally, we cut out the lenses with circular punchers aftercuring the silicone at room temperature overnight. We have developedprocesses and techniques to build sensors as low as 50 μm thick so thatoverall contact lens sensor can be less than 150 μm.

For the auxetic sensor version, the only difference in manufacturing wasin the step 4 of FIG. 13, in which we have used a patterned film insteadof a flat film as the bottom layer. The patterning was done as shown inFIG. 15.

Variations and Modifications

-   1) The microfluidic strain sensing principles could be used for wide    range of medical applications where strain sensing is necessary.    Biomedical applications other than glaucoma management could be    listed as; physiotherapy monitoring (e.g. at joints in hand    injuries), speech recognition, fetus/baby monitoring, tremor    diseases, robotics etc.-   2) Microfluidic strain sensing can be used for biosensing and    biochemical sensing. For example, it can be used to monitor to    measure the strain applied by cells on a surface. Mechanical cues    play important role in cellular processing such as cell    differentiation, apoptosis, and motility. Cells senses and exerts    forces on substrate that they grow. Tumor cells generate more forces    than regular cells. Shear stress, one of the leading physical cues    is causing upregulation of genes activated by mechanic signals.    Understanding mechanical cues generated by cells will be crucial to    understand cancer progression which is triggered by mutations of    mechanotransduction pathways of cells. Our strain sensor will    provide direct monitoring of direct cancer cells signaling under    exposure of different physical and mechanical cues. Therefore, it    will bring novel approach in cancer studies. By using our sensor,    new biomarkers will be discovered as well as new drug therapies    could be implemented. These devices will also help in several other    conditions including regulation of synaptic plasticity of neurons as    forces are one of the key factors for progress of synaptic    plasticity.

To understand cell's response to different conditions. Two layers ofmicrofluidic channels can be built as shown in FIG. 16. As cells grow,we could image the strain sensor on the bottom channel. This willprovide tissue stiffening. Top channel can also be manipulated byapplying different flow rate which changes the shear stress. In thisdesign, cells mechanical response can be observed while they are beingmechanically manipulated. This design will be used in biomarker and drugdevelopment.

Cancer tissues as they progress shows more stiffer character. Inaverage, cancer cell will have 4 times stiffer than regular tissues.Understanding earlier stiffness of cancer cells will lead to earliercancer detection. The strain sensor could be incorporated into patcheswhich can be externally used on the skin. Specifically, it could be usedin skin and breast cancer types. Such patches with infrared beadsembedded in microchannel could be optimized and implanted to internalorgans in the case of ovarian cancer, liver and brain cancers.Especially, these patches could be implanted after severe tumor removalsurgeries to monitor cancer reoccurrence. Combining microfluidics-basedstrain sensors with flexible silicon electronics will enable multiplexedmeasurements on three dimensional soft tissues in vivo. This signalcould be transferred to cloud-based system using wi-fi embeddedtechnologies. Overall, the strain sensors incorporated with advanceelectronics will provide continuous monitoring of tissues which carrieshigh chance of cancer reoccurrence.

-   3) The miLenS can either be manufactured by: i) embedding strain    sensor with the desired shape/size in to a contact lens, as    described or ii) directly patterning the desired topographies on the    surface of the contact lens through soft lithography where features    on a mold transferred to the contact lens.-   4) The distance between the microscopic geometric features on the    contact lens can be directly measured instead of using    microfluidics. This distance will change as a function of IOP. The    geometric shapes and patterns of these features should be carefully    selected to maximize the sensitivity to IOP. The IOP will be    measured based on the imaging of contact lens sensor with    geometrical features (geoLenS) similar to miLenS. FIG. 17 shows the    top and side views of the example geoLenS. The location and shapes    of the microscopic features to be used for IOP determination are    shown. Besides the star shape shown in the top view and side view,    other example shapes are also provided. The combinations of these    shapes can also be used. In the top view, the radius of the contact    lens is denoted by r and the value of r can be between 0.5 and 1 cm.    θ, shows the angle between the features positioned at the periphery    of the contact lens and it determines the number of features that    will be placed angularly on a contact lens. θ could be in between    10° (36 features at the periphery) and 180° (Two features at the    periphery). Minimum of two features needed on the contact lens. d₁,    d₂, d₃, . . . d_(n) denote the distances between consecutive    features and can be between 0.01 to 1 cm. The total distance    d=d₁+d₂+d₃+ . . . +d_(n) should be smaller than r. The radius of    curvature of the contact lens, r_(c), shown in side view can be    between 0.5 to 1 cm. The characteristic width of features, w could    be 0.001 to 0.5 cm.

As the IOP changes, the distances between peripheral features, e.g., d₁,change and can be used as a measure of the IOP change. The distancesbetween central features, e.g., d₂ or d₃, or the width of any feature,w, can be used as a reference measurement because they do not change inresponse to IOP. The distance between the opposing features at theperiphery (total distance is 2 d) changes the most as response to IOPchange. The distance of any one of the geoLenS features to the knownfeatures of the eye (i.e. iris border) can be detected as a measure ofIOP.

To test the feasibility of the concept proposed above, we fabricated acontact lens, which was made of PDMS and has thickness ˜250 μm. Fortesting, we fabricated a realistic eye model made of PDMS as shown inU.S. Provisional Patent Application 62/556,366 filed Sep. 9, 2017 (FIG.19-left therein), which is incorporated herein by reference.

The radius of curvature of the eye model changes ˜4 μm/mmHg (3 μm/mbar)and this is very close to the behavior of human's eye.

We put marks on the contact lens and we placed it on the eye model wecreated as shown in U.S. Provisional Patent Application 62/556,366 filedSep. 9, 2017 (FIG. 19-right therein), which is incorporated herein byreference. These marks served as probes and enabled us to measure thechange in distance between different locations on the contact lens as afunction of applied pressure. We had four levels of applied pressure inthe eye model varying from 25 mbar to 100 mbar. We sampled fourlocations on the contact lens (3 distance measurements) and distancesbetween these locations are plotted as a function of applied pressure asshown in U.S. Provisional Patent Application 62/556,366 filed Sep. 9,2017 (FIG. 29 therein), which is incorporated herein by reference. Thepoint located on the center of the contact lens is labeled as location‘1’ and the number is increased as the points located further from thecenter (e.g., location ‘2’). The distances between different markedpoints (e.g., location ‘1’ to location ‘2’) were measured. In FIG. 20,blue, red, and green lines show the distance as a function of appliedpressure for location 1 to 2, location 2 to 4, and location 4 to 6,respectively. Corresponding linear fits are plotted as well. Overall,the preliminary results show that the distances between differentlocations on the geoLenS follow a linear function of applied pressureand this is in a measurable range.

-   5) The geoLenS features can be fabricated similar to miLenS or they    can just be marked with an ink.-   6) The miLenS reservoir channels can have a serpentine shape instead    of circular.-   7) The device can be used as a temperature sensor as it is sensitive    to thermal expansion of the material.-   8) The device is insensitive to air pressure changes. It can be used    in vacuum, e.g. in space applications.-   9) The images can be taken by a smartphone camera, a special    handheld camera, or by a wearable camera. The images can be taken    directly across the eye, at 45° angle or at 90° angle or any angle    between 0°-90° angle.-   10) The front or back camera of the smartphone can be used for    imaging.-   11) The images can be collected by the patient, at will or    automatically when the patient is reading something on the phone.-   12) The image analysis can be made by the microprocessor of the    camera or can be transferred to a main server for further    processing.-   13) The patient can pay for subscription to cloud services such as    data storage, analysis etc.-   14) The miLenS channels can be filled with a colored liquid to    improve the contrast on the iris or sclera.

Additional Technical Notes

The invention pertains to a closed microfluidic network for strainsensing applications. The device has strain sensitivity of 2-15 mminterface movement per 1% strain depending on the number of rings. Thesensitivity can be increased even further by increasing the number ofrings. It is robust enough to withstand pressure changes that areapplied for 24 hours and has a lifetime of months. These features makeit attractive for applications where extremely strain levels smallerthan 0.1% need to be measured for time periods longer than 2 hours. Wehave embedded the sensor into a contact lens for monitoring intraocularpressure (IOP). The required detection limit for IOP is 1 mmHg. Thiscorresponds to a strain of 0.05%. We have achieved this strain detectionlimit by designing a liquid reservoir network which includes multiplemicrofluidic channels as a liquid reservoir. The liquid reservoirnetwork is connected to a sensing channel and the sensing channel isconnected to an air reservoir. These three components form a closedsystem. The sensor with its three components in one possibleconfiguration is shown in FIG. 3. FIG. 3 is the top view of the sensorshowing when it is embedded in a contact lens. The sensor is filled fromthe inlet with a working liquid, using only capillary forces. When theworking liquid reaches the outlet, both inlet and outlet are sealedusing the sensor material to form a closed system with a fixed liquidvolume inside. At this point the liquid fills the sensing channel,approximately half of its total length, creating a liquid/air interface.Both the contact lens and the sensor are made of elastomers such assilicone and polyurethane but can be made of other materials such assilicone/hydrogel.

The sensor works based on volume amplification of microfluidic liquidreservoir network when it is stretched under tangential forces. Theworking principle of the sensor is described in FIG. 4. Here, anotherconfiguration of the sensor components, where they are linearlydistributed instead of radially distributed, is used for simplicity. Theside view of the sensor with a liquid reservoir which could havemultiple chambers, A), versus a single wide chamber, B), is compared.When the sensor is stretched under tangential forces, the shape of thesensor and of its components change as depicted in A*) and B*),respectively. 410-A and 410-B are representations of the possible stressregions on the sensor in the vicinity of liquid reservoir. Forreference, the total initial length of the sensor is shown as 1-1′,total initial liquid reservoir network width is shown as 2-2′, andinitial position of the liquid air interface is shown as 3. There arethree notable mechanical changes which could occur when such a closedmicrofluidic network is stretched under tangential force;

i) Elongation: When A*) and B*) are compared with A) and B),respectively, it can be seen that the total sensor length, (1-1′), willincrease due to elongation. Similarly, the liquid reservoir networkwidth, (2-2′), will also increase.

ii) Collapses: In the case of single reservoir, the thin membrane abovethe liquid reservoir will collapse due to the induced stress and due tothe low rigidity of this membrane, as shown in B*). When multiplechambers with higher rigidity membranes are used, the collapses will notoccur, or will decrease significantly, as shown in A*).

iii) Liquid reservoir volume increase and resulting vacuum effect: Whenthe liquid reservoir width is elongated, its total volume will increaseif the membrane collapses can be prevented or reduced significantly.This volume increase can be amplified if the liquid reservoir consistsof multiple chambers with small widths as shown in B*). Theamplification will be even higher when auxetic patterns are created onthe membrane of the small reservoir chambers. When the volume of theliquid reservoir increases, this causes a vacuum effect and this vacuumpulls the liquid/air interface position (3) towards the liquidreservoir. The movement of this interface, in μm, per IOP change, inmmHg, is defined as sensitivity. Each 1 mmHg IOP change causes a strainof 0.05% according to literature. This strain causes approximately 100μm position change on the interface position.

Another factor that should be considered for maximum sensitivity is theYoung's modulus (E) of the sensor material. Increasing the E reduces thecomfort. When contact lenses with high lubricity is used for improvedcomfort, the contact friction between the cornea and sensor/lens willdecrease, which will cause slipping and decreased sensitivity,especially for high E sensors. According to our experimental andsimulation results, the optimal E is in the range of 0.2-10 MPa formaximum sensitivity and comfort. As the E is reduced below 2 MPa, thewidth of the reservoir channels also has to be reduced below 100 μm.

The invention claimed is:
 1. A microfluidic strain sensing device formonitoring a intraocular pressure change of an eyes, the microfluidicstrain sensing device comprising: (a) a contact lens; (b) a closedmicrofluidic network embedded with the contact lens, wherein the closedmicrofluidic network has a volume sensitive to an applied strain andwherein the closed microfluidic network further comprises: (i) a gasreservoir containing a gas, (ii) a liquid reservoir containing a liquidthat changes volume when the strain is applied, and (iii) a sensingchannel able to hold the liquid within the sensing channel, wherein thesensing channel connects to the gas reservoir on one end and connects tothe liquid reservoir on another end, wherein the sensing channelestablishes a liquid-gas equilibrium pressure interface and equilibriumwithin the sensing channel, which fluidically change as a response toradius of curvature variations on a cornea, or as a response tomechanical stretching and release of the cornea, and wherein theliquid-gas equilibrium pressure interface and equilibrium are used formeasuring the intraocular pressure.
 2. The intraocular pressuremonitoring device as set forth in claim 1, wherein the liquid reservoirforms at least one ring and wherein the gas reservoir is positionedinside the at least one ring.
 3. The intraocular pressure monitoringdevice as set forth in claim 1, wherein the liquid reservoir volume issensitive to a tangential force on the eye relative to a radial force onthe eye wearing the contact lens.
 4. The intraocular pressure monitoringdevice as set forth in claim 1, wherein the liquid reservoir has astiffness in a radial direction and/or a smaller channel width relativeto a stiffness in a tangential direction.
 5. The intraocular pressuremonitoring device as set forth in claim 1, wherein the contact lens is asilicone contact lens, a hydrogel contact lens or a combination thereof.6. The intraocular pressure monitoring device as set forth in claim 1,wherein the sensing channel has a strain sensitivity of about 4.5 mminterface movement per about 1% strain applied to the liquid reservoir.7. The intraocular pressure monitoring device as set forth in claim 1,wherein the sensing channel has an inner diameter of about 1-10 mm. 8.The intraocular pressure monitoring device as set forth in claim 1,wherein the sensing channel has an inner diameter of about 5-12 mm witha cross sectional area of 10-11-10-8 m².
 9. The intraocular pressuremonitoring device as set forth in claim 1, wherein the liquid reservoirhas one or more chambers.
 10. The intraocular pressure monitoring deviceas set forth in claim 1, wherein the liquid reservoir has one or morechambers with concentric rings.
 11. The intraocular pressure monitoringdevice as set forth in claim 1, wherein the liquid reservoir has one ormore chambers with concentric rings, wherein the concentric rings areconnected at one or more locations.
 12. The intraocular pressuremonitoring device as set forth in claim 1, wherein the liquid reservoirhas a surface with a pattern.
 13. The intraocular pressure monitoringdevice as set forth in claim 1, wherein the liquid reservoir has aceiling and a floor, the ceiling having a convex shape and wherein theceiling is curved towards the floor.
 14. The intraocular pressuremonitoring device as set forth in claim 1, wherein the contact lens hasno actively controlled components or electrical components.
 15. Theintraocular pressure monitoring device as set forth in claim 1, whereinthe closed microfluidic network is transparent.
 16. The intraocularpressure monitoring device as set forth in claim 1, wherein the closedmicrofluidic network is oleophobic.